Examinations are carried out with the aid of computed tomographs in many problem areas in medicine. In this case, the computed tomograph includes an X-ray source and a detector module, opposite the X-ray source, with a number of measuring channels that are formed by individual detector elements. For the purpose of spatially resolved detection of the X-radiation, the detector elements are generally arranged next to one another in one or more rows. Findings relating to the distribution of material within the object to be examined can be obtained from the spatially resolved measurement of the attenuation of the X-radiation by the object to be examined located between the X-ray source and detector module.
Known on the one hand, for detecting X-radiation are detectors having indirect converters that are assembled from a scintillator material with a downstream photodetector. The scintillator converts the incident X-radiation into optical radiation that is subsequently detected by the photodetector. The number of photons produced per X-ray quantum is generally approximately proportional to its quantum energy in this case. This technique uses integration over the electric signal received by the photodetector over a prescribed time interval. The intensity of the received X-radiation is then yielded by dividing the value integrated by the detector by the mean quantum energy, to be determined stochastically, per X-ray quantum.
Also known for detecting X-radiation are specific semiconductor materials in which the incident X-radiation generates charge carriers directly. The number of the charge carriers generated in these direct converters per X-ray quantum is generally approximately proportional to its quantum energy in this case.
A counting method is also known for computed tomographs instead of integration over the analog electric signal received by the converter. Thus, DE 102 12 638 A1 exhibits a detector module for a computed tomograph that has a number of detectors or measuring channels which detect the X-radiation on the basis of direct converters. Each converter is connected to a pulse generator for generating counting pulses as a function of the received electric signals. The pulse generator relays the pulses to a counting device that counts the received counting pulses over a prescribable time period and outputs the result.
In this configuration, the electronics of the detectors have substantially fewer analog parts than the electronics of conventional detectors. The electronics provided can therefore be smaller, more cost-effective and more noise-immune. The detector module represented in this printed publication further includes a threshold logic unit composed of a number of parallel-connected comparators as part of the pulse generator, each of the comparators being assigned a counter in the counting device in order to be able to count X-ray quanta of different energy independently of one another. The detection of the incident X-radiation is enabled in this way with regard both to intensity and to the quantum energy of the individual X-ray quanta.
In addition to computed tomography (CT), positron emission tomography (PET) has also become increasingly widespread in medical diagnostics of recent years. Whereas computed tomography is concerned with an anatomical imaging technique, PET permits, for example, the visualization and quantification of metabolic activities in vivo.
Positron emission tomography utilizes the particular properties of positron radiators and positron annihilation for the purpose of quantitative determination of the functioning of organs or cell zones. The patient is administered appropriate positron radiators in this case before the examination. During decay of a positron radiator, a proton is converted into a positron, a neutron and a neutrino. However, the positron is not directly detected, since its range is limited to a few mm. In the patient's tissue, the positron is braked by scattering processes at the shells of neighboring atoms and is captured by a shell electron.
Annihilation produces two gamma quanta that fly apart in opposite directions. The energy of the two gamma quanta is 511 keV in each case on the basis of the law of conservation of energy and impulse. If the gamma quanta of two opposite detector elements is measured within a specific time, the location of the annihilation is established at a position on the connecting line between these two detector elements. This is utilized for generating images in PET.
A combined measurement using both CT and PET techniques is desired in many instances on the basis of the different information the two techniques return. Thus, for example, the article by T. Beyer et al., entitled PET/CT-Tomographie mit neuem PET-Detektormaterial für ultraschnelle Bildgebung in der klinischen Onkologie [PET/CT tomography with the aid of new PET detector material for ultrafast imaging in clinical oncology], electromedia 70 (2002), pages 167–172 discloses a combined PET/CT tomograph which can be used to produce complementary PET and CT images in a short time in a single examination. The detector modules for computed tomography and for positron emission tomography are mounted in the same installation in this case such that images produced with the aid of the two techniques can be registered exactly and without any problem. US 2003/0004405 A1 also discloses a combined PET and X-ray CT tomograph that can be used to take CT and PET pictures of an object to be examined directly one after another. The two detector modules, the detector module for CT and the detector module for PET, are mounted in this case on a common support inside the gantry.
The two abovementioned printed publications indicate that so far CT detectors and PET detectors have been implemented using different technologies. The detector modules for CT systems are generally based on integrating detectors, for example having GdOS or CdWO scintillators. By contrast, PET systems operate with counting scintillator detectors based on BGO and LSO.
At first glance, the measuring requirements differ greatly between the systems. Computed tomography must process very large quantum currents, the time resolution being in the region of from 200 μs to 600 μs in this case. By contrast, PET makes use of a coincidence measurement. This requires a high time resolution which is, for example, approximately 300 ns given the use of BGO and approximately 30–50 ns given the use of LSO, while the quantum fluxes are smaller by orders of magnitude than in the case of computed tomography. Instead of an X-ray tube, a positron emission tomograph makes use in the body of the patient of the decaying radionuclides at functional key groups, for example at tumor cells, as radiation source.
As already explained, upon decay of the nuclides, two gamma quanta are emitted in opposite directions. In the case of PET the detector module must generally cover the greater part of the gantry arc length for the purpose of detection. It is subdivided into detector elements having a side length of a few mm. Each detector element, also designated as a measuring channel in the present patent application, generates upon detection of a gamma quantum an event record that specifies the time and the detection location, that is to say the corresponding detector element.
These information items are transmitted to a fast logic unit and compared. If two events coincidence within a maximum time period, it is assumed that there is a gamma decay process on the connecting line between the two associated detector elements. As in the case of CT, the PET image is then reconstructed using a tomography algorithm, that is to say the so called back projection.
U.S. Pat. No. 6,449,331 B1 describes a combined PET and CT detector module that is mounted on a support inside the gantry of a combined PET and CT installation. The individual detector elements of the detector module respectively include a scintillator crystal with a downstream photodetector. LSO is used as scintillator material, in the way known from positron emission tomographs.
The photodetector is connected, on the one hand, to an event detector that registers all the reception events, and, on the other hand, to an integration unit that integrates over the received signal. The detector module operates in three operating modes.
In a first operating mode as PET detector, it outputs an item of information relating to the time and the location of a reception event via the event detector. In a second operating mode, the detector module operates as a standard CT detector in the case of which the signal received by the photodetector is integrated with the aid of the integration unit. In a third operating mode, individual events are counted in the CT operation in order to obtain a CT image therefrom.
In the course of the dead times of the indirect converter (scintillator crystal and photodetector), this third operating mode can be used only given very small quantum fluxes of the X-radiation, and so it is generally necessary to switch over to this second operating mode in the integration unit given the quantum fluxes customary for CT. The implementation of such a detector module remains, however, complicated since, in addition to the event detector there is also a need for the integration unit for CT measurements.